1. Field of the Invention
The present invention relates generally to gas transfer systems and more specifically to an extracorporeal blood oxygenator utilizing microporous fibers.
2. Description of the Prior Art
In the late 1970s, technology was developed to allow the extrusion of very thin hollow tubes from polymeric materials. Additional benefits of this technology allowed for the creation of a multiplicity of micropores within the wall of such tubing. Such microporous hollow fibers (MHF) were soon readily adapted for use in extracorporeal blood oxygenators, serving as the membranous element separating blood from the gaseous phase, in order to minimize the blood trauma experienced when blood is directly mixed with gases.
Predating the development of MHF, blood oxygenators had been constructed so as to directly mix gaseous oxygen with the blood to achieve the necessary gas transfer. As previously stated, this resulted in significant blood trauma and necessitated the subsequent defoaming of the blood/gas mixture prior to return of the blood to the patient's arterial circulation. A significant incidence of patient pathology, such as gaseous embolization and silicone defoaming agent embolization to the brain and other major organs, when utilizing such bubble oxygenators led early investigators to develop a potentially safer device called a membrane oxygenator.
Early membrane oxygenator designs utilized flat sheets of membranous material (usually thin silicone rubber sheets) to separate the alternating blood and gas channels of the device. The membrane material was arranged in either a flat stack, referred to as a flat plate oxygenator or as a continuous coil around a central core, referred to as a spiral coil oxygenator. Blood was allowed to flow within the channel between two opposing membrane layers, while gaseous oxygen flowed within the adjacent channels on the other side of the membrane sheets. The necessary transfer of oxygen molecules into the blood and the simultaneous removal of carbon dioxide molecules from the blood was by passive diffusion and was limited by the chemical solubility of the gases in the membrane material. While many of these early devices were functional, and were utilized clinically, they were relatively inefficient gas transfer devices requiring large membrane surface areas to provide adequate gas exchange. An additional problem with the earlier designs was the necessary compression and/or membrane tension required to keep the membranes from bulging apart during active blood flow. Such bulging resulted in undesirably thick blood film thickness within the device and caused further deterioration of gas transfer efficiency and increased fluid priming volume requirements.
With the development of microporous polymeric materials, a hybrid design was possible combining the gas transfer efficiency benefits of a direct blood gas interface (as in the bubble oxygenator) with the reduced blood trauma benefits of a membrane device. The molecules of gas transiting between phases within the device no longer had to physically dissolve within the membrane material in order to pass between channels, as gas molecules could directly pass through the fluid/gas film interface created within the micropores of the membrane. Passage of gross gaseous emboli into the blood channels was prevented, due to the high surface tension of the plasma fluid film at the surface of the microscopic pores of the membrane material. Early use of microporous polymeric sheets in flat plate designs encountered the same problems of control of blood film thickness as previously mentioned. The emergence of MHF, however, allowed more consistent control of this previously difficult variable in membrane oxygenator design, by providing a fixed dimension lumen within the fiber for the passage of blood while simultaneously bathing the exterior of the MHF with oxygen gas. Such early MHF membrane oxygenators utilized a linear bundle of fibers, the ends of which were first sealed within a block of polymeric material followed by the reopening of each end of the fiber lumens by cleanly slicing off the distal edge of the cured resin blocks. In this manner, direct communication of the gas and blood passages was prevented at the terminal ends of the fiber bundle.
The early luminal blood flow MHF membrane oxygenators, generally referred to as "internal flow configuration oxygenators," were eventually supplanted by development of "external flow configuration oxygenators" in which the blood flow passed over the external surface of the MHF while gaseous oxygen was allowed to flow within the internal lumen of the MHF. Such a design change evolved due to the high fluid pressures which occurred as a result of passing relatively viscous fluid through extremely narrow fiber lumens within the "internal flow configuration oxygenators". Initial "external flow configuration oxygenator" designs utilized the same linear fiber bundle, requiring various techniques of fiber bundle compression to maintain the smallest blood film possible around each fiber. Subsequently it was found that by controlled tension spiral winding of the MHF around a central core, a tubular core of MHF could be created which more closely controlled the blood film thickness, or boundary layer. Controlling the boundary layer in this manner allowed the adult membrane oxygenator total fiber surface area to be reduced from approximately 4.5 square meters to 2.0 square meters due to boundary improvement in gas exchange efficiency. The spiral wound membrane oxygenator is currently the most popular design, due to its efficiency, in spite of significant manufacturing difficulties with the uniform creation of the spiral wound fiber cores.
Recently, other methods for control of boundary layer and external flow configuration MHF membrane oxygenators have appeared in the market place. One such method is the bundling of MHF fibers into "fiber ribbons" by tightly wrapping a small fiber bundle with a retaining thread. These bundles are then placed within a channel in a metallic coil, which serves both as a blood flow channel and a thermal exchange surface.
A second functional feature required for the successful use of a membrane oxygenator in cardiopulmonary support of the cardiac surgical patient is that of heat exchange. Provision for adequate thermal exchange within the cardiopulmonary bypass circuit must be made in order to maintain and/or alter the patient's body temperature during the surgical procedure. This is most commonly achieved by utilizing a blood oxygenator which has an integral heat exchanger as part of its design, although this function may also be accomplished by inclusion of a separate heat exchanger somewhere within the extracorporeal circuit. With early bubble oxygenators, the integral heat exchangers were placed on the outflow, or arterial, side of the oxygenator such that heat exchange occurred after the gas exchange process had been completed. It was found, however, that when warming the blood in such configurations micro bubbles of gas could be detected, due to the decreased solubility of gases in a fluid as temperature of the solution is increased. Accordingly, subsequent designs provided for heat exchanger placement on the inlet, or venous, side of the oxygenator.
It is of interest to note that all currently utilized oxygenators have either venous or arterial side heat exchangers, with the exception of the aforedescribed device utilizing "fiber ribbons" in a metallic coil. The metallic blood channels, within which the fiber ribbons are placed, serve as a thermal exchange surface, as the undersurface of the metallic coil is fitted with water conduits. By passing thermally conditioned water through these conduits, the metallic coil can be either heated or cooled, to achieve blood heat exchange within the same channels utilized for gas exchange.
It is important in any blood oxygenator that it provide an efficient system for transferring gas to and from the circulating blood. It is also of critical importance that the device be capable of cooling the blood being recirculated into the patient's vascular system so that the patient's body temperature can be cooled to produce a physiologically protective hypothermic state. It is conversely important that the device be capable of warming the blood so that near the end of a surgical procedure, the device can warm the recirculating blood that is returning to the patient to a normothermic state.
Another important feature of a membrane oxygenator is that it has a minimal fluid priming volume. The cardiopulmonary bypass circuit is normally composed of numerous components, with the oxygenator and other components interconnected by significant lengths of sterile tubing. Additional lengths of tubing are connected to the patient's vascular system and are utilized to direct the patient's venous blood into the extracorporeal circuit, and to return the arterialized blood to the patient's arterial circulation. This circuit must be completely filled with an appropriate physiologic fluid, prior to connection into the patient's vascular system to prevent catastrophic embolization of gas into the circulatory system of the patient. Obviously, the larger the total fluid volume of the bypass circuit, the greater the hemodilutional effect on the patient. As one progressively dilutes the patient's blood, a critical point will be reached at which the patient's blood will not be able to transport sufficient oxygen to support tissue requirements without excessive blood flow rates. Such extreme hemodilution will then require transfusion of homologous blood into the circuit to increase the blood's oxygen-carrying capacity. Consequently, the optimal design of an oxygenator would be to minimize the fluid priming volume required for safe operation.
Still another important feature of a blood oxygenator is that it exert minimal trauma on the blood. Blood trauma can occur in many different ways with the most significant cause being that of excessive sheer forces acting upon the blood elements flowing through the device. In the optimal design for a membrane oxygenator, one must balance the need for an extremely thin boundary layer of blood next to the membrane (to maximize gas transfer) with the need to keep blood velocity minimized (to reduce sheer force induced blood trauma).
Obviously, in light of the fact that blood oxygenators are used in critical surgery such as cardiopulmonary bypass, it is extremely important that the device perform predictably and reliably. It is also important due to the increasing costs of medical care that the cost of manufacturing be minimal.
Several patents have been issued for devices developed to address individually or in combination some of the issues raised hereinabove. By way of example, U.S. Pat. No. 4,791,054 to Hamada, et al., U.S. Pat. No. 4,111,659 issued to Bowley, U.S. Pat. No. 3,998,593 issued to Yoshida, et al. and U.S. Pat. No. 5,137,531 issued to Lee, et al. all concern oxygenator type devices wherein the oxygenator has separate and distinct chambers for oxygenation and temperature control.
U.S. Pat. No. 3,342,729 issued to Strand discloses a permeability separatory cell that utilizes a mesh membrane of fibers having cation exchange properties running in one direction and fibers having anion exchange properties in a perpendicular direction. U.S. Pat. No. 3,794,468 issued to Leonard discloses a mass transfer device having a wound tubular diffusion membrane. U.S. Pat. No. 4,722,829 issued to Giter is another illustration of a blood oxygenator wherein tubes through which treating gas flows include spherical lobes which are interengaged to define small passages through which the blood moves. Finally, U.S. Pat. No. 4,940,617 issued to Baurmeister discloses a multilayered hollow fiber wound body wherein the fibers are wound either in helices or spirals.
It is to address the issues identified above and to resolve the issues in a more satisfactory manner than the existing prior art that the present invention has been developed.